1. Field of the Invention
The present invention generally relates to nuclear medicine, and systems for obtaining nuclear medicine images of a patient's body organs of interest. In particular, the present invention relates to a novel detector configuration for nuclear medical imaging systems that are capable of performing either positron emission tomography (PET) or planar and single photon emission computed tomography (SPECT).
2. Description of the Background Art
Nuclear medicine is a unique medical specialty wherein radiation is used to acquire images that show the function and anatomy of organs, bones or tissues of the body. Radiopharmaceuticals are introduced into the body, either by injection or ingestion, and are attracted to specific organs, bones or tissues of interest. Such radiopharmaceuticals produce gamma photon emissions that emanate from the body. One or more detectors are used to detect the emitted gamma photons, and the information collected from the detector(s) is processed to calculate the position of origin of the emitted photon from the source (i.e., the body organ or tissue under study). The accumulation of a large number of emitted gamma positions allows an image of the organ or tissue under study to be displayed.
Two basic types of imaging techniques are PET or “coincidence” imaging, and single photon imaging, also known as planar or SPECT imaging. PET imaging is fundamentally different from single photon imaging. In PET, events are detected from the decay or annihilation of a positron. When a positron is annihilated within a subject, two 511 KeV gamma rays are simultaneously produced which travel in approximately opposite (i.e., 180°) directions. Two scintillation detectors are positioned on opposite sides of the patient such that each detector will produce an electrical pulse in response to the interaction of the respective gamma rays with a scintillation crystal. In order to distinguish the detected positron annihilation events from background radiation or random events, the events must be coincident (i.e., both occur within a narrow time window) in each detector in order to be counted as “true” events. When a true event is detected, the line connecting the positions of the two points of detection is assumed to pass through the point of annihilation of the positron within the subject being imaged.
By contrast, single photon imaging, either planar or SPECT, relies on the use of a collimator placed in front of a scintillation crystal or solid state detector, to allow only gamma rays aligned with the holes of the collimator to pass through to the detector, thus inferring the line on which the gamma emission is assumed to have occurred. Both PET and single photon imaging techniques require gamma ray detectors that calculate and store both the position of the detected gamma ray and its energy.
Present day single photon imaging systems all use large area scintillation detectors (on the order of 2000 cm2). Such detectors are made either of sodium iodide crystals doped with thallium (NaI(Tl)), or cesium iodide (CsI). Scintillations within the NaI crystal caused by absorption of a gamma photon within the crystal, result in the emission of a number of light photons from the crystal. The scintillations are detected by an array of photomultiplier tubes (PMTs) in close optical coupling to the crystal surface. Energy information is obtained by summing the signals from the PMTs that detected scintillation photons, and position information is obtained by applying a positioning algorithm to the quantitative signals produced by the PMT array. The original gamma-ray camera is described in U.S. Pat. No. 3,011,057 issued to Hal Anger in 1961.
The CsI camera is typically used with either a single silicon-based photodiode detector or an array of silicon-based photodiode detectors, which detect scintillation events emitted from the CsI crystal. CsI crystals are used where the relatively low cost, ruggedness and spectral response of the CsI crystal are desired in favor of alternative crystal materials such as NaI.
In PET imaging, scintillation crystals with short response times are required in order to properly detect the coincidence events with high timing resolution. Typically such crystals are chosen from among materials such as NaI, BGO, LSO and BaF2. Detectors coupled to such crystals can be an array of PMTs, a single “position-sensitive” PMT (“PS-PMT”), or fast-response silicon-based photodiodes such as avalanche photodiodes.
Because the conventional Anger camera uses a thin planar sheet or disk of scintillation crystal material, it is necessary to cover the entire field of view of the crystal with light detectors such as PMTs or photodiodes. Additionally, the sampling capability of such scintillation crystals could be improved by increasing the number of gamma photons emanating from an imaging subject that are absorbed by the crystal, and consequently increasing the number of scintillation events that can be detected for use in constructing an image.
The bar detector is a specific configuration of scintillation detector that has been used in astronomical and high energy physics applications. The bar detector consists of an elongated scintillation crystal bar having a relatively small cross section. A photosensor such as a PMT is optically coupled to each end of the bar. The light from a gamma photon event within the scintillation crystal volume is detected by the two PMTs. The timing or signal information can be used to determine the location of the event in the bar. Additional bars can be placed next to each other for two-dimensional detection.
An example of a proposed design for a PET detector module using a bar detector is given in Moses et al., “Design Studies for a PET Detector Module Using a PIN Photodiode to Measure Depth of Interaction,” IEEE Transactions on Nuclear Science NS-41, pp. 1441-1445 (1994), incorporated herein by reference in its entirety. According to this design, a scintillation bar is coupled at one end to a PMT, and at the other end to a photodiode, in order to measure the depth of interaction (DOI) of the scintillation event in the bar.
In past bar detector experiments for physics and astronomy, NaI (and sometimes CsI) bars of up to 100 cm were used to detect gamma photons of up to 10 MeV. Positional resolution within the bar ranged from 1.5 cm at 200 keV to 2 cm at 10 MeV, although worse resolutions were reported. An energy resolution of 9.4% and a timing resolution of 10 ns at 662 keV and a 100 cm NaI bar were reported by a physics group for a balloon borne gamma telescope. Energy resolutions from other experiments were higher for the same energy gamma photon. These studies have cited geometry, bar size, light attenuation coefficient and electronic noise as the major factors in determining the spatial and energy resolution of bar detectors. However, the performance of bar detectors as designed in the prior art is insufficient for use in medical imaging applications.
It would be desirable to use a bar detector in medical imaging applications. First, the bar detector can achieve the same field of view as a conventional Anger camera with a significantly smaller number of PMTs. This is because a scintillation event occurring anywhere within the length of the bar can be detected by the light sensors optically coupled to the ends of the bars using a DOI calculation. Thus, photodetectors do not need to be placed over the entire field of view as in the conventional Anger camera. Additionally, multiple bars can be placed next to each other in a two dimensional array, creating a detector module than can localize the gamma photon interaction in a three dimensional volume, allowing for depth of interaction (DOI) decoding capability. This DOI information can enhance the spatial resolution of a positron emission tomograph system by improving sampling characteristics of the detector system, as the total thickness of the module can be larger than the thickness of a single crystal and thus can interact with a higher percentage of gamma photons emanating from a subject. Additionally, for SPECT systems, the bar detector module presents a very economical modality for constructing a camera with substantial advantages over a conventional planar camera (see FIGS. 3a-3b), including high throughput and elimination of the necessity for gantry motion.